1. Field of the Invention
The present invention is directed to a real-time method for imaging organic tissue. The method employs a polarization-enhanced fluorescence imaging system in order to obtain images of the organic tissue. The obtained images can be used in the demarcation of, among other things, nonmelanoma skin cancers.
2. Background of the Related Art
Nonmelanoma skin cancers are the most common forms of human cancer. About 75% of all skin cancers are basal cell carcinomas (BCC) and about 20% are squamous cell carcinomas (SCC). These cancers are a major cause of morbidity in the Caucasian population. They commonly appear on sun-exposed areas of the body such as the head and neck. Since many tumors occur on the face it is imperative to preserve normal skin surrounding the tumor. Unfortunately, most of these tumors have poorly defined boundaries, which makes visual detection of the tumor borders and, consequently, precise excision a challenging problem.
In the US, Mohs micrographic surgery (MMS) is an accepted procedure that removes as little normal skin as possible while providing the highest cure rate. Using detailed mapping and complete microscopic control of the excised lesion the Mohs surgeon can pinpoint areas at the surgical margins involved with cancer that are otherwise invisible to the naked eye. While precise and accurate, MMS is also a time-consuming and staff-intensive procedure. It requires a surgeon trained in dermatopathology, a dedicated laboratory and a technician to prepare and evaluate frozen sections. Because of these shortcomings, MMS is used in the minority of cases.
In recent years, the development of optical imaging modalities led to the introduction of techniques that may become a viable alternative to the existing methods of skin tumor detection and demarcation. However, all these techniques lack one or more elements necessary for their practical use in a clinical setting.
Confocal reflectance microscopy has been used to study different normal and pathological skin conditions. It allows imaging within turbid media with high resolution (lateral ˜1 μm, axial (section thickness) ˜3 μm), which is comparable to histology. The major disadvantage of in vivo confocal microscopy for the assessment of skin tumor margins is the small field of view (0.25 mm to 0.3 mm). By sacrificing axial resolution (˜30 μm) it is possible to enlarge the field of view up to 2 mm. But even 2 mm field of view is much smaller than the size of most lesions. To examine the entire suspected cancerous area the sequences of images should be captured and stitched together. This process takes time and the resulting image may be distorted by patient's motion.
Polarized light has been used extensively for biological and medical applications. In most cases, skin cancer arises from epidermis. As such, for detection of skin lesions, it is advantageous to acquire superficial images. To achieve this goal while retaining a large field of view polarized light imaging may be employed. The use of polarized light allows imaging of the superficial tissue layers only. When the light incident on the sample is linearly polarized, the two images acquired with the co-polarized and cross-polarized light can be used to largely isolate the single-scattered component, which arises mainly from superficial skin layers. If a turbid medium, like skin, is illuminated with a linearly polarized light, the backscattered light partially retains its polarization. The light that is specularly reflected and single-scattered by a random turbid medium has the same polarization as the incident beam.
For example, let IMps be the intensity of single scattered light at the image plane. Multiple scattering randomizes polarization of the propagating beam. Eventually, half of the multiple backscattered light has the same polarization (IMpm), and another half has a polarization transversal (IMsm) to the incident beam polarization, consequently IMpm=IMsm.
Conventional image, IM, is created by single scattered light and multiple scattered light:IM=IMps+IMpm+IMsm.  (1)
An image acquired using the remitted light polarized in the direction parallel to the polarization of the incident light is created by a sum of single scattered light and multiple scattered light:IMp=IMps+IMpm.  (2)
An image acquired using the remitted light polarized in the direction transversal to the polarization of the incident light is created by multiple scattered light only:IMs=IMsm.  (3)
The difference image, DIM, is obtained by subtraction:DIM=IMp−IMs=IMps+IMpm−IMsm=IMps.  (4)
This image is formed by single scattered light since, as it was explained above, IMpm=IMsm.
Single backscattering happens in skin, depending on the wavelength of light, pigmentation, and blood content, at the depth of approximately 70 μm to 200 μm in the visible and near infrared spectral range. Recently, white polarized light digital imaging was being used to evaluate pigmented skin lesions. (Jacques S L, Roman J R, Lee K: Imaging superficial tissues with polarized light. Las. Surg. Med. 2000; 26:119-129.) A polarization image, PIM, was created and analyzed:
                    PIM        =                                            IM              p                        -                          IM              s                                                          IM              p                        +                          IM              s                                                          (        5        )            
The numerator is equal to DIM. The denominator is a conventional image. The ratio of the difference image to the conventional image can be used to cancel out the contrast that is associated with any superficial chromophore (i.e. melanin, blood) present in the tissue. The thickness of the imaged layer is about 200 μM (white light). Melanin strongly scatters light, producing bright areas with excellent contrast in pigmented lesions. Such high contrast based on scattering would not be expected to occur reliably in nonmelanoma skin cancers, which contain variable amounts of melanin. Thus, this method includes an inability to use spectral information encoded in white light image for lesion characterization and comparatively poor contrast of the nonmelanoma cancer lesion in the image.
Considerable efforts have been devoted to the development of skin tumor imaging techniques based on detection of endogenous fluorescence and exogenous fluorescence of photosensitizers. In Brancaleon et al. the possibility of using autofluorescence (endogenous fluorescence) spectroscopy for the detection of nonmelanoma skin cancer was explored. (Brancaleon L, Durkin A J, Tu J H, Menaker G, Fallon J D, Kollias N: In vivo fluorescence spectroscopy of nonmelanoma skin cancer. Photochem. Photobiol. 73(2): 178-183, 2001.) Their in vivo and in vitro studies have shown that the endogenous fluorescence of tryptophan residues was stronger and fluorescence associated with collagen and elastin was weaker in tumor than in normal tissue. At the same time the authors mentioned that in the case of morpheaform BCC, when collagen fibers are surrounded with tumor cells, and SCC in situ, when there is no tumor invasion into the dermis, the collagen fluorescence might increase. The loss of collagen and elastin fluorescence in the vicinity of a tumor was observed for 78% of fresh frozen cancer tissue samples. The areas characterized by the loss of fluorescence were two- to threefold larger than the tumor size determined from histological evaluation. Therefore, the method suggested in the paper may be applied for nonmelanoma cancer detection, but cannot be used for precise tumor demarcation during surgery.
Photodynamic therapy (PDT) has also been tried as an alternative method for treatment of skin cancers. An example of a PDT procedure for dermatology involves the topical application of δ-aminolevulinic acid (ALA) followed by irradiation with red light (λ˜635 nm). ALA is a precursor in the biosynthesis of protoporphyrin IX (Pp IX) that accumulates in tumor tissue. When cells containing Pp IX are irradiated with red light, they are selectively killed. Pp IX is fluorescent, and therefore, may be used for tumor detection.
Wennberg et al imaged in vivo the areas of Pp IX fluorescence and compared the location and the size of these areas with the size of the lesions determined by histological methods. (Wennberg A M, Gudmundson F, Stenquist B, Ternesten A, Moelne L, Rosen A, Larko O: In vivo detection of basal cell carcinoma using imaging spectroscopy. Acta Derm. Venereol. 79: 54-61, 1999.) They found that in 50% of lesions the correlation with histology was good, in 23% the correlation was partial, and in 27% there was no correlation at all. The authors noticed that the selectivity of Pp IX fluorescence is not high enough, since in several cases Pp IX fluorescence was detected from sun-damaged skin, healing scars, and normal hair follicles. Similar studies, which employed multi-wavelength fluorescence and lifetime fluorescence imaging, were conducted by several other groups. The predictive capability of Pp IX fluorescence imaging and its ability to demarcate lateral extent of the tumor are still questionable. (Hewett J, Nadeau V, Ferguson J, Moseley H, Ibbotson S, Allen J W, Sibbett W, Padgett M: The application of a compact multispectral imaging system with integrated excitation source to in vivo monitoring of fluorescence during topical photodynamic therapy of superficial skin cancers. Photochem. Photobiol. 73(3): 278-282, 2001; Andersson-Engels S, Canti G, Cueddu R, Eker C, af Klinteberg C, Pifferi A, Svanberg K, Svanberg S, Taroni P, Valentini G, Wang I: Preliminary evaluation of two fluorescence imaging methods for the detection and the delineation of basal cell carcinomas of the skin. Las. Surg. Med. 26:76-82, 2000.)
In many cases the differences of optical signals from normal and diseased tissues are subtle, therefore a lot of effort is devoted to the development and evaluation of novel optical contrast agents. Gold nanoparticles and microspheres filled with light scattering media are examples of such contrast agents. (West J L and Halas N J: Applications of Nanotechnology to Biotechnology—Commentary. Current Opinion in Biotechnology 11: 215-220, 2000.) The advantage of these contrast agents is the tunability of their optical properties. In other words, using the structure, the size, and the refractive index as variable parameters it is possible to create the particles, which will enhance scattering and/or absorption of the tissue containing these particles at the specific predefined wavelengths.
The utility of such contrast agents greatly depends on the efficiency of their delivery to the target tumor tissue. One of the approaches is to design very selective contrast agents that could be injected intravenously and that would migrate and localize into a tumor. Another one is to bind the existing contrast agents to specific cell surface proteins thus achieving tumor selectivity. (Bugaj J E, Achilefu S, Dorshow R B, Rajagopalan R. Novel fluorescent contrast agents for optical imaging of in vivo tumors based on a receptor-targeted contrast agent-peptide conjugate platform. J Biomed Opt 6:122-33, 2001.)
The development of these state of the art molecular specific contrast agents is a complex and challenging problem. Elaborate and time-consuming animal model testing is required to evaluate the potential of these approaches for in vivo tumor imaging in humans. At this point of development it is not feasible to attempt application of these experimental contrast agents in clinical practice. Where BCC is concerned, investigation of applicability of such molecular-specific agents is even more problematic, since there exists no animal model for BCC of human skin. Therefore an approach based on utilization of the existing contrast agents appears to be more suitable for applications in present clinical settings.
The nontoxic contrast agents that are selectively retained by cancerous tissue have been applied previously to aid in visual examination of oral, bladder, and cervix lesions. Phenothiazinium contrast agents including methylene blue (MB) and toluidine blue (TB) in particular have been used for staining various carcinomas in vivo. (Kaisary A V: Assessment of radiotherapy in invasive bladder carcinoma using in vivo methylene blue staining technique. Urology, 28(2): 100-102, 1986; Eisen G M, Montgomery E A, Azumi N, Hatmann D-P, Bhargava P, Lippman M, Benjamin S B: Qualitative mapping of Barrett's metaplasia: a prerequisite for intervention trials. Gastrointestinal Endoscopy, 50 (6): 814-818, 1999.) Phenothiazinium contrast agents are accumulated to a much greater extent in mitochondria of carcinoma cells compared to normal cells. (Oseroff A R, Ohuoha D, Ara G, McAuliffe D, Foley J, Cincotta L: Intramitochondrial contrast agents allow selective in vitro photolysis of carcinoma cells. Proc. Natl. Acad. Sci. USA, 83: 9729-9733, 1986.) MB has been successfully applied to grossly demarcate neoplastic tumors in bladder, tumors of pancreas, and Barrett's esophagus metaplasia. TB has been used topically to detect oral carcinoma, and Barrett's esophagus metaplasia.
TB is a preferred stain for use in Mohs surgery for BCC, because TB staining provides some advantages relative to hematoxylin-eosin (H&E) including a highly identifiable staining pattern (metachromasia) of BCC. Since TB is routinely used to stain fresh-frozen tissue sections during MMS, the processed polarized light images of stained tumors are remarkably similar to standard Mohs micrographic surgery maps. This similarity significantly simplifies the process of image understanding and interpretation for a Mohs surgeon. Hair follicles, sebaceous glands, fat, and normal stromal elements are visible in detail, and appear differently from the tumor, which appears very dark due to the increased, relative to the normal tissue, uptake of the contrast agent.
Another type of contrast agents used is fluorescent contrast agents (fluorophores). Fluorescent contrast agents absorb light at the specific wavelengths and emit light at longer wavelengths. There are several fluorophores that preferentially stain tumors and, therefore, can be used for tumor detection and demarcation. These include: Pp IX, tetracycline, TB, and MB. Fluorescence spectra are sensitive to the changes in biochemical environment of the fluorophore molecules. Biochemical composition of the diseased skin differs significantly from the normal. Upon excitation with polarized light, the emission from fluorescent samples is also polarized (Lakowic J R: Principles of Fluorescence Spectroscopy: NY, Plenum Press, 1983, Feofilov P P, lrv. Akad Nauk SSSR. Ser. Fiz. 9, 317, 1945, Chen R F and Bowman R L, 1965). This polarization is a result of the photoselection of fluorophores, according to their orientation relative to the direction of the polarized excitation. Fluorescence polarization is defined as P=(I∥−I⊥)/(I∥+I⊥).
Alternatively, fluorescence anisotropy, which is defined as Ir=(I∥−I⊥)/(I∥+2 I⊥) can be evaluated. To measure fluorescence polarization, the sample is excited with linearly polarized light. When the observing polarizer is oriented parallel (∥) to the direction of the polarized excitation the observed image (or intensity) is I∥. When the polarizer is oriented perpendicular (⊥) to the polarization plane of the excitation light the observed image is I⊥.
Fluorescence emission can be depolarized by a number of phenomena, including rotational diffusion of the fluorophore during the lifetime of the excited state, energy transfer, reabsorption, etc. The dependence of the fluorescence polarization on rotational diffusion led to numerous applications of this technique in biochemical research. For example, fluorescence polarization measurements have been used to study intracellular structural changes, quantify protein denaturation and rotational rates of proteins. There also exist several clinical applications of fluorescence polarization assays. These assays are used for therapeutic drug monitoring, for determination of fetal lung maturity etc. Therefore fluorescence polarization imaging gives a promise of further increasing the specificity of the contrast agents like TB, MB, and TCN.
All of the above mentioned approaches attempted so far for bedside imaging of skin cancer have advantages and drawbacks. It appears, however, that all these techniques lack one or more elements necessary for their practical use in a clinical setting. Thus, confocal microscopy, although providing a superior spatial resolution, suffers from extremely limited field of view and complexity of implementing multi-wavelength imaging.
In addition, this technique is very sensitive to small changes in the position of the investigated object. White-light polarization imaging, being simple and inexpensive, at the same time is unable to use spectral information and as a result does not provide necessary contrast. PpIX-fluorescence imaging is not specific enough; whereas the autofluorescence imaging tends to exaggerate tumor dimensions and is not capable of localizing the morpheaform BCC.
Thus, in view of the above-described deficiencies, a simpler, more accurate and time-efficient method would be desirable for mapping tumor borders.